1. Field of the Invention
This invention relates to gamma or positron emission tomography (PET) cameras. This invention also relates to an improved light distribution method for reducing the costs associated with manufacturing the crystal arrays for the PET cameras.
2. Background
A PET camera typically consists of a polygonal or circular ring of radiation detection sensors 10 placed around a patient area 11, as shown in FIG. 1. Radiation detection begins by injecting isotopes with short half-lives into a patient's body placeable within patient area 11. The isotopes are absorbed by target areas within the body, causing the isotope to emit positrons that are detected when they generate gamma rays. When in the human body, the positrons collide with electrons and the two annihilate each other, releasing gamma rays. The emitted rays move in opposite directions, leave the body and strike the ring of radiation detectors 10.
The ring of detectors 10 includes an inner ring of scintillation crystals 12 and an outer ring of light detectors or photomultiplier tubes 14, shown in FIG. 2. The scintillation crystals respond to the incidence of gamma rays by emitting a flash of photon energy (scintillation) that is then converted into electronic signals by a corresponding adjacent photomultiplier tube 14. A computer records the location of each energy flash and then plots the source of radiation within the patient's body by comparing flashes and looking for pairs of flashes that arise from the same positron-electron annihilation point. It then translates that data into a PET scan image. The PET monitor displays the concentration of isotopes in various colors indicating level of activity. The resulting PET scan image indicates a transaxial view of neoplasms or tumors existing in the patient's body.
Early PET scanners required a single photomultiplier tube to be coupled to a single scintillation crystal. A crystal can be made very narrow (e.g., 1 mm). The thinner the crystal, the greater the resolution of the PET camera. However, narrow crystals are useful only if the crystal location can be accurately decoded. The smallest available photomultiplier tube (PMT) is somewhat large by comparison (e.g., 10 mm). Hence, a practical advance in PET scanners allows a single PMT to service several crystals. As disclosed in my U.S. Pat. Nos. 4,733,083 and 4,883,966, a single PMT can service several crystals. Because PMTs are relatively expensive (a single PMT may alone cost between $250 and $650), minimizing the number of PMTs can drastically reduce the cost of the PET camera.
As shown in FIG. 3, two PMTs 14a and 14b can service a row of eight scintillation crystals 12 in a known design. The eight crystals are formed from a unitary crystal block 16. As described in U.S. Pat. No. 4,743,764, each crystal 12 is formed by placing slots or cuts 18 at varying depths into block 16. The depth of each cut determines the amount of photon energy being directed to a respective PMT 14a or 14b. For example, crystal 12a is formed having a cut 18 placed the entire depth of block 16 and separating crystal 12a from crystal 12b. Photon energy generated within crystal 12a is directed entirely into the right-side PMT 14a. Typically, cut 18 is filled with light reflecting materials or the sides of the cut are polished so as to effectively prevent photon energy from passing across cuts. Shortening the depth of cut 18 will allow photon energy to be directed along the shortened cut distance and then disperse past the cut edge. For example, photon energy within crystal 12b will disperse toward the center of the block to both PMT 14a and 14b. Photon energy will predominantly strike PMT 14a; however, some energy will strike PMT 14b due to the absence of the cut extending the entire depth of the left hand side of crystal 12b.
The layout of scintillation crystals 12 is in three dimensions. FIG. 3 illustrates the x and z axes of eight crystals within a single block 16. However, FIG. 4 illustrates the x and y axes of four blocks 16, each block adjacent four PMTs 14. Typically an array of crystals are formed between cuts within a block. The blocks are then joined side-by-side to preferably form a ring surrounding the patient area. In two dimensions (x and y), a single PMT 14 can, for example, service sixteen crystals 12 (i.e., a 4.times.4 array of crystals 12) as shown in FIG. 4. Likewise, four PMTs can service a block or 8.times.8 array of crystals.
Shown in FIGS. 2-4 are various conventional crystal/PMT arrangements, which typically involve placement of the outer edge of a PMT adjacent to and aligned with the outer edge of an array of crystals 12 (or edge of a block 16). For purposes of isolating the specific crystal being scintillated, it is important that the photon energy sent to the respective PMT identify where, within the array of crystals serviced by the PMT, that the scintillation occurred. Scintillation occurring within one crystal is often directed to one of many PMTs associated with the entire block of crystals. In FIG. 4, for example, if crystal H is scintillated, then all photon energy will be directed to PMT 14b due to the cut being the entire depth of the crystal block as shown in FIG. 3. If crystal N is scintillated, then photon energy will be primarily directed to PMT 14b; however, PMT 14a, 14e and 14f will also receive photon energy. If the location of scintillation is nearer the center of the array (e.g., crystal C' is scintillated), then PMT 14b will read only slightly more photon energy than PMTs 14a, 14e and 14f. By determining the relative amounts of photon energy received by four PMTs associated with a single array or block, the relative X and Y location of scintillation can be determined within that array or block, the resolution of the X and Y location being directly proportional to the width of each crystal 12. Of course, it is implicit that the system must be capable of decoding the crystal locations.
The scintillation light from one crystal is distributed to four PMTs for decoding the position from that crystal by comparing the ratio of scintillation signal received by the four PMTs as follows: ##EQU1## where a, b, e, and f are the amount of scintillation signal received from each of the four PMTs 14a, 14b, 14e, and 14f in FIG. 4. For example, when the corner crystal A, detects a gamma ray, both the X-position and Y-position will be near maximum (=1.0) because almost all of the scintillation light goes to PMT 14a and very little light will be distributed to the other three PMTs 14b, 14e, and 14f. Another example is when crystal B' detects a gamma ray. The scintillation light will be equally distributed to all four PMTs (14a, 14b, 14e, and 14f); the X-position will be approximately equal to 0.5 and the Y-position will also be close to 0.5.
Unfortunately, with the outer edge of the array being aligned with and adjacent to the outer edge of the PMTs (i.e., two sides of each PMT), there is an upper limit on the size of the PMT in relation to the size of the array. If, for example, two (as opposed to four) PMTs service a single array, then predominate photon energy in one PMT indicates that scintillation occurred at the left (or upper) half of the array as opposed to the right (or lower) half. Thus, only single axis (e.g., X axis) detection is possible. Single axis detection is not sufficient to identify with precision the exact location of scintillation within the block or array. Rather, a double-axis detection scheme is needed whereby both the X and Y location of each scintillation is detectable. If, for example, a single PMT services a single array, no indicia whatsoever of the location scintillation with the array can be determined. The PMT cannot differentiate where, within the serviced array, the scintillation occurred. Thus, as long as conventional detection schemes utilize "non-offset" PMT/array edges, the number of PMTs per array cannot be effectively decreased.